The present invention relates generally to magnetic resonance imaging (“MRI”) systems and methods and, more particularly, the invention relates to systems and methods for magnetic resonance angiography (“MRA”) and perfusion imaging.
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B0), the individual magnetic moments of the nuclei in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a magnetic field (excitation field B1) that is in the x-y plane and that is near the Larmor frequency, the net aligned moment, Mz, may be rotated, or “tipped”, into the x-y plane to produce a net transverse magnetic moment Mxy. A signal is emitted by the excited nuclei or “spins”, after the excitation signal B1 is terminated, and this signal may be received and processed to form an image.
When utilizing these “MR” signals to produce images, magnetic field gradients (Gx, Gy, and Gz) are employed. Typically, the region to be imaged is scanned by a sequence of measurement cycles in which these gradients vary according to the particular localization method being used. The resulting set of received MR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques.
The measurement cycle used to acquire each MR signal is performed under the direction of a pulse sequence produced by a pulse sequencer. Clinically available MRI systems store a library of such pulse sequences that can be prescribed to meet the needs of many different clinical applications. Research MRI systems include a library of clinically-proven pulse sequences and they also enable the development of new pulse sequences.
The MR signals acquired with an MRI system are signal samples of the subject of the examination in Fourier space, or what is often referred to in the art as “k-space.” Each MR measurement cycle, or pulse sequence, typically samples a portion of k-space along a sampling trajectory characteristic of that pulse sequence. Many pulse sequences sample k-space in a raster scan-like pattern sometimes referred to as a “spin-warp,” a “Fourier,” a “rectilinear,” or a “Cartesian” scan. The spin-warp scan technique employs a variable amplitude phase encoding magnetic field gradient pulse prior to the acquisition of MR spin-echo signals to phase encode spatial information in the direction of this gradient. In a two-dimensional implementation (“2DFT”), for example, spatial information is encoded in one direction by applying a phase encoding gradient, Gy, along that direction, and then a spin-echo signal is acquired in the presence of a readout magnetic field gradient, Gx, in a direction orthogonal to the phase encoding direction. The readout gradient present during the spin-echo acquisition encodes spatial information in the orthogonal direction. In a typical 2DFT pulse sequence, the magnitude of the phase encoding gradient pulse, Gy, is incremented, ΔGy, in the sequence of measurement cycles, or “views” that are acquired during the scan to produce a set of k-space MR data from which an entire image can be reconstructed.
There are many other k-space sampling patterns used by MRI systems. These include “radial,” or “projection reconstruction” scans in which k-space is sampled as a set of radial sampling trajectories extending from the center of k-space. The pulse sequences for a radial scan are characterized by the lack of a phase encoding gradient and the presence of a readout gradient that changes direction from one pulse sequence view to the next. There are also many k-space sampling methods that are closely related to the radial scan and that sample along a curved k-space sampling trajectory rather than the straight line radial trajectory.
An image is reconstructed from the acquired k-space data by transforming the k-space data set to an image space data set. There are many different methods for performing this task and the method used is often determined by the technique used to acquire the k-space data. With a Cartesian grid of k-space data that results from a 2D or 3D spin-warp acquisition, for example, the most common reconstruction method used is an inverse Fourier transformation (“2DFT” or “3DFT”) along each of the 2 or 3 axes of the data set. With a radial k-space data set and its variations, the most common reconstruction method includes “regridding” the k-space samples to create a Cartesian grid of k-space samples and then performing a 2DFT or 3DFT on the regridded k-space data set. In the alternative, a radial k-space data set can also be transformed to Radon space by performing a 1 DFT of each radial projection view and then transforming the Radon space data set to image space by performing a filtered backprojection.
Magnetic resonance angiography (“MRA”) uses the magnetic resonance phenomenon to produce images of the human vasculature. To enhance the diagnostic capability of MRA a contrast agent such as gadolinium can be injected into the patient prior to the MRA scan. Typically, one of the tricks with this contrast enhanced (“CE”) MRA method is to acquire the central k-space views at the moment the bolus of contrast agent is flowing through the vasculature of interest. Collection of the central lines of k-space during peak arterial enhancement is key to the success of a CE-MRA exam. If the central lines of k-space are acquired prior to the arrival of contrast, severe image artifacts can limit the diagnostic information in the image. Alternatively, arterial images acquired after the passage of the peak arterial contrast are sometimes obscured by the enhancement of veins. In many anatomic regions, such as the carotid or renal arteries, the separation between arterial and venous enhancement can be as short as 6 seconds.
The short separation time between arterial and venous enhancement dictates the use of acquisition sequences of either low spatial resolution or very short repetition times (“TR”). Short TR acquisition sequences severely limit the signal-to-noise ratio (“SNR”) of the acquired images relative to those exams in which longer TRs are possible. The rapid acquisitions required by first pass CE-MRA methods thus impose an upper limit on either spatial or temporal resolution.
As indicated above, the acquisition of MRA data is timed such that the central region of k-space is acquired as the bolus of contrast agent arrives in the arteries of interest. The ability to time the arrival of contrast varies considerably and it is helpful in many applications where proper timing is difficult to acquire a series of MRA image frames in a dynamic study that depicts the separate enhancement of arteries and veins. Such temporal series of image frames is also useful for observing delayed vessel filling patterns caused by disease. This requirement has been partially addressed by acquiring a series of time resolved images using a 3D “Fourier” acquisition. When a dynamic study is performed the time resolution of the study is determined by how fast the k-space data can be acquired for each image frame. This time resolution objective is often compromised in order to acquire all the k-space data needed to produce image frames of a prescribed resolution without undersampling artifacts.
Perfusion imaging is employed to assess the viability of tissues. An exemplary perfusion imaging method includes administering a contrast agent to the subject, after which a series of MR images are acquired as the contrast agent perfuses into the tissues of interest. From this series of contrast-enhanced MR images hemodynamic parameters such as blood flow, blood volume and mean transit time may be computed.
Hemodynamically weighted MR perfusion images of cerebral blood flow (“CBF”) may be acquired and used in combination with diffusion-weighted (“DWI”) MR images to delineate regions of viable brain parenchyma that are at risk of further infarction. The DWI MR image shows ischemic regions where brain cells have died, and the CBF image shows regions with reduced blood flow that indicates at risk tissue. The size of the “ischemic penumbra” surrounding ischemic tissues is a critical component in evaluating treatment options.
It is possible to assess regional cerebral hemodynamics by analyzing MR signal intensity changes after the first pass of the paramagnetic contrast medium. While passing through the capillary network, a short bolus of contrast material produces local magnetic field inhomogeneities that lead to a reduction in the transverse magnetization relaxation time T2*of the bulk tissue. This susceptibility effect can be recorded by a series of rapid T2*-weighted gradient echo images that reveal how the MR signal changes during the first pass of the contrast agent. The resulting MR signal intensity versus time curves can be converted into contrast agent concentration-time curves. By using the indicator dilution theory, two important hemodynamic parameters can be determined from these curves: the CBF, known as tissue perfusion, and the cerebral blood volume (“CBV”). However, the concentration of contrast agent in the arterial blood pool, the so-called “arterial input function” (“AIF”), must be known if absolute quantification of the CBV and CBF measurements are to be achieved. Typical methods used to measure the AIF require a step in which the operator manually selects a region of interest (“ROI”), based on anatomic information, that depicts an artery. The concentration-time curve from all voxels included in the ROI is then used to calculate the AIF.
Given the clinical usefulness of both MRA images and MR perfusion images, some have attempted to combine the acquisition of the information for both image types. For example, some have combined dynamic contrast-enhanced (DCE) MRI perfusion images with MRA images in a serial fashion. Of course, as addressed above, CE-MRA imaging presents the need to time the passage of the peak arterial contrast and maintain a separation between arterial and venous enhancement. In a serial perfusion and MRA acquisition, performing perfusion imaging before the CE-MRA data acquisition leads to unwanted venous contamination in the subsequent time-resolved MRA image. On the other hand, performing the time-resolved MRA before perfusion imaging confounds the subsequent perfusion study because the baseline background signal is enhanced.
Accordingly, some have proposed interleaving segments of a 3D MRA data acquisition with multiple complete 2D perfusion image data acquisitions. To make such a combination of two separate pulse sequence more tolerable, some proposed collecting few phase-encoding lines and shortening the TR of the 2D perfusion data acquisitions. However, doing so lowers the overall data acquisition time by reducing the resolution of the acquired perfusion images. Similarly, attempts to sacrifice the temporal resolution of the perfusion and/or MRA images to control the duration of the overall data acquisition time are limited, for example, at least by the speed and timing of the contrast enhancement. In any case, even when clinical needs can tolerate substantial sacrifices in spatial or temporal resolution of the resulting images, these interleaving methods inevitably increase the acquisition times and can present substantial challenges in coordinating contrast passage during imaging acquisition and tolerating less-than-ideal contrast enhancement.
Therefore it would be desirable to have a system and method for acquiring angiographic and perfusion images using MRI in a coordinated fashion that does not unduly extend acquisition times or require particular data to be acquired during periods timed to undesirable phases of contrast enhancement for the data to be acquired.